Modern medical imaging techniques have important applications in health care. Modalities such as X-ray computed tomography (CT), magnetic resonance imaging (MRI) and ultrasound imaging are the main tomographic techniques available in most modern medical centers. Visible-light endoscopy is another major imaging modality which is used extensively in procedures like bronchoscopy or colonoscopy. Each of these techniques employs different physical principles and measures different properties of the biological tissue under study With different resolution. Further, they can commonly be performed in-vivo. A third type of imaging, optical microscopy, is still utilized widely in clinical medicine. However, optical microscopy is currently limited to detailed examination of excised or resected specimens and is not used in-vivo. In many circumstances, the superior contrast and resolution afforded by optical microscopy is such that physical biopsy followed by optical microscopic histology is considered the gold standard for diagnosis.
Combinations of these techniques, such as using a low-resolution tomographic modality along with high-resolution imaging, biopsies or interventional procedures are constantly being studied and evaluated. Evaluation of these techniques is based on technological feasibility, clinical benefit and cost.
Optical Coherence Tomography (OCT) is a relatively new imaging technique based on the low-coherence property of electromagnetic radiation that enables high-resolution depth profilometry in a turbid, highly scattering media such as biological tissue. Its use in biomedical imaging is currently being investigated in several research and industrial laboratories. The main advantage of OCT lies in its ability to localize the depth of reflection from a sub-surface site in tissue. This localization is essentially determined by the coherence properties of the light source used and can be as low as 2 to 20 μm for selected near-IR sources (e.g. lasers or amplified spontaneous emission devices). This gives a measure of the depth resolution attainable with OCT. Independently of the coherence characteristics, the lateral resolution is determined by the beam cross-section at the depth of imaging and by the lateral spacing of the acquired data. Typical values for lateral spacing in the literature are in the 5 to 30 μm range. The price to be paid for this remarkable cross-sectional imaging ability in intact turbid tissue is the limited imaging depth since, due to multiple scattering and absorption, both coherence and penetration of light are degraded resulting in OCT imaging depths of approximately 2 to 3 mm.
Most current implementations of OCT are based on Michelson interferometry with a 50/50 beam splitter directing the incident coherent light beam into a reference path containing a mirror (i.e. a reference arm) and a sample path containing the interrogated sample (i.e. a sample arm). Both free-space optic and fiber-optic implementations of this scheme are currently used. Reflected beams from the mirror in the reference arm and from the tissue in the sample arm are recombined in the same splitter and half of the resultant light energy impinges on a detector. Incoherent superposition of the two light fluxes typically occurs except when the optical path lengths of the two beams are matched to within the coherence length of the source. Within this limited distance, the coherent superposition of the two light fluxes yields an interference pattern with a fringe magnitude that is proportional to the reflectivity of the tissue at that particular depth. Depth profiling of the sample is then achieved by scanning the reference arm length or more correctly by scanning the optical path length of the reference arm by using a time delay in the reference arm (this is equivalent to lengthening the reference arm). Various detection methods to measures and quantify these faint amplitude modulations amidst large background diffuse reflectance have been developed having a dynamic range of approximately 70 to 110 dB. Furthermore, lateral translation of the beam and axial motion of the reference mirror enables one to construct a two-dimensional reflectivity picture over a desired field of view. Means of improving the final image quality, such as performing image processing through de-convolution, have also been investigated.
The foregoing is a brief description of conventional reflectivity OCT imaging. Other variations include, for example, flow (Doppler) imaging and polarization imaging (albeit at the expense of additional, complexity of the OCT optics and/or signal processing techniques). Images from these additional techniques are usually obtained in conjunction with images from conventional OCT so some image overlay or fusion is possible. Upon further technological development and/or clinical implementation, may add sufficient information content to increase the clinical utility of OCT in medicine.
However, many OCT designs and approaches that have been successfully implemented in tabletop research systems are not directly suitable for in-vivo imaging such as in gastroenterologic or bronchoscopic endoscopy. Instead, they may be more suitable for dermatological, ophthalmologic and dental applications. In contrast, in-vivo OCT imaging must address the issues of speed, resolution, contrast, penetration and instrument size. Images must be obtained sufficiently quickly to negate the effects of patient motion while still achieving suitable axial and lateral resolution; and maintaining an instrument size which is small enough to be endoscopically useful. Powerful near-IR sources, fast means of altering the reference arm length and custom-designed distal optical devices have been successfully developed to overcome the difficult challenges posed by in-vivo endoscopy.
The latest OCT technology employs a single-mode optical fiber with distal side-viewing optics introduced into the accessory channel of a conventional white-light endoscope. To build-up an image, the viewing direction of the OCT fiber is either linearly scanned to and fro over an approximate 2 mm distance, or is rotated via a flexible guide-wire or interlocking gear mechanism at several revolutions per second. Simultaneous with this translation or rotation, the reference arm length outside the endoscope is rapidly varied via an optical phase delay to generate depth scans (i.e. A-scans). Currently, these OCT systems operate at frame rates up to conventional video rates but more typically at 4 to 8 frames per second, with a frame presenting a fully circumferential view to a depth of 2 to 3 mm. The resultant resolution values are approximately 5 to 25 μm in the depth (axial) direction and approximately 20 to 40 μm in the lateral direction. As well, the lateral resolution generally degrades with an increase in distance from the fiber tip of the OCT device due to geometric divergence. These OCT systems have a dynamic range which is somewhat lower than that of corresponding ex-vivo systems due to increased noise levels and faster imaging speeds.
Based on the latest OCT technology, it is questionable whether coherent in-vivo OCT systems are adequate for successful clinical imaging. The images are certainly useful, but substantial improvement is required if the elusive goal of “optical biopsy” is to be realized. For example axial and lateral resolution can be improved. While the improvement in the former usually involves the use of better low-coherence sources (i.e. CW and pulsed sources), the issue of sub-optimal and depth-varying lateral resolution is more difficult to address. In ex-vivo systems, with its relaxed constraints of speed and physical size, lateral resolution is improved by focusing the beam to a few microns with a high-NA (numerical aperture) objective lens. In contrast to conventional OCT scanning, the imaging can now be performed in the lateral (en face) direction with a pre-selected depth with small oscillations in path length difference, followed by a small depth increment as necessary. The high-NA objective lens is often coupled to tissue via a refractive-index matching liquid. The general approach of using OCT with a high-NA distal optic lens is known as Optical Coherence Microscopy (OCM). However, the improved lateral resolution at the beam waist location comes at the expense of lateral blurring at other depths because the highly focused beam has a very shallow depth of field. Thus, the lens-to-surface distance must be varied to focus to different depths. In addition, a dynamic tracking scheme is needed to keep the location of the coherence gate (within which coherent interference between the optical beams from the sample arm and the reference arm is possible) and the beam waist at the same depth. These techniques for lateral resolution improvement have not been attempted during in-vivo endoscopy because of size and speed requirements.